The present invention relates to implantable medical devices, and more particularly to an implantable cardioverter defibrillator (ICD) configured to provide a high efficiency defibrillation waveform.
An ICD continues to be a relatively large device for implantation in the human body. The size of the ICD is primarily determined by the battery and capacitors used therein. The size of the battery (or batteries, in some instances) and capacitors, in turn, is determined by the shock energy requirements for a defibrillation pulse. Thus, a design approach that reduces the energy requirements for defibrillation results in a direct reduction in the overall ICD size.
In existing ICD devices, the defibrillation waveform or pulse used to deliver a defibrillation shock to the heart is generated by first charging the equivalent of a single capacitor (most ICDs use two capacitors connected in series to function as a single capacitor, thereby reducing the working voltage requirements for each capacitor of the series stack, as explained below) to a desired charge level (voltage) and then discharging the single capacitor through the cardiac tissue for a prescribed period of time during a first or positive phase of the defibrillation waveform, and then reversing the polarity of the discharge for a second prescribed period of time during a second or negative phase of the defibrillation waveform, thereby producing a biphasic stimulation pulse or waveform. It should be noted that in this context the term xe2x80x9csingle capacitorxe2x80x9d is used to refer to a single capacitance, which may be, and usually is obtained by a hardwired connection of two capacitors in series such that the two series capacitors always function and act as though they were a single capacitor. (Two or more capacitors are connected in series in this manner in order to achieve a higher working voltage for the series-connected capacitor. That is, when two capacitors are connected in series, and each has a working voltage of, e.g., 375 volts (V), then the overall or total working voltage of the series combination becomes 750 V.)
The purpose of applying a defibrillation shock to the heart is to shock the heart out of a state of fibrillation, or other non-functional state, into a functional state where it may operate efficiently as a pump to pump blood through the body. To this end, the positive phase of the biphasic waveform is preferably a very high voltage that serves to synchronously capture as many heart membrane cells as possible. See, Kroll, xe2x80x9cA minimum model of the signal capacitor biphasic waveformxe2x80x9d Pace, November 1994. The negative phase of the biphasic waveform, in contrast, simply serves to remove the residual electrical charge from the membrane cells and bring the collective membrane voltage back to its original position or value. See, e.g., Kroll, supra; Walcott, et al., xe2x80x9cChoosing The Optimal Monophasic and Biphasic Wave-Forms for Ventricular Defibrillation,xe2x80x9d, Journal of Cardiovascular Electrophysiology (September 1995). A biphasic pulse generator of the type used in an ICD device is shown, e.g., in U.S. Pat. Nos. 4,850,357, issued to Bach, Jr.; and U.S. Pat. No. 5,083,562, issued to de Coriolis et al.
When a voltage shock is first applied to a membrane cell, the membrane does not respond to the shock immediately. Rather, the cell response lags behind the applied voltage. This time lag is more or less predictable in accordance with the Blair membrane model. See, e.g., Blair, xe2x80x9cOn the intensity-time relations for stimulation by electric currents. Ixe2x80x9d J. Gen Physiol., Vol. 15, pp. 709-729 (1932), and Blair, xe2x80x9cOn the intensity time relations for stimulation by electric currents. IIxe2x80x9d, J. Gen Physiol., Vol. 15, pp. 731-755 (1932); Pearce et al., xe2x80x9cMyocardial stimulation with ultrashort duration current pulses,xe2x80x9d PACE, Vol. 5, pp. 52-58 (1982). When the applied voltage comprises a biphasic pulse having a constant voltage level for the duration of the positive phase (a condition achievable only when the voltage originates from an ideal battery), the membrane cell response to the positive phase reaches a peak (i.e., is at an optimum level) at the trailing edge of the positive phase. Unfortunately, when the applied voltage originates from a charged capacitor, as is the case for an ICD device, the applied voltage waveform does not remain at a constant voltage level, but rather has a significant xe2x80x9ctiltxe2x80x9d or discharge slope associated therewith. Such tilt or slope causes the peak membrane cell response to occur at some point prior to the trailing edge of the positive phase, which is less than optimum. What is needed, therefore, is a way to optimize the applied voltage waveform so that a maximum membrane cell response occurs coincident with, or nearly coincident with, the trailing edge of the positive phase.
It is known in the art to switch the capacitors of an ICD from a parallel configuration during the positive phase of a biphasic defibrillation pulse to a series configuration during the negative phase of the biphasic defibrillation pulse. See, e.g., U.S. Pat. No. 5,199,429 (FIG. 7A) and U.S. Pat. No. 5,411,525. While such action produces a defibrillation waveform having a somewhat different shape, i.e., a waveform having a leading edge voltage of the second or negative phase which is approximately twice the trailing edge voltage of the first or positive phase, such action does little to achieve a maximum cell membrane response coincident with the trailing edge of the first or positive phase.
It is also known in the art to sequentially switch capacitors in an ICD device in order to allow waveform xe2x80x9ctailoringxe2x80x9d, e.g., prolong the positive phase duration by sequentially switching in a second charged capacitor as shown in FIG. 6A of U.S. Pat. No. 5,199,429, or by sequentially switching in second, third and fourth charged capacitors, as shown in FIG. 6C of U.S. Pat. No. 5,199,429. However, such xe2x80x9ctailoringxe2x80x9d still does not address the main concern of achieving a maximum cell membrane response coincident with the trailing edge of the positive phase.
It is thus evident that what is needed is a capacitor switching scheme and/or method for use within an ICD device which achieves a maximum cell membrane response near or coincident with the trailing edge of the positive phase.
It is also desirable to provide an ICD that is as small as possible. The limiting factor on ICD thickness is the diameter of the high-energy capacitors. As indicated above, current ICDs typically use two electrolytic capacitors. Current technology in electrolytic capacitors limits the stored voltage to about 450 V per capacitor. Therefore, the current approach is to use two large (200 xcexcF or more) capacitors to achieve the stored energy of 25 J-40 J required for defibrillation. Therefore, the thickness of the ICD is determined by the thickness of the large capacitors. There is thus a need for an ICD construction, which would permit the needed energy for defibrillation to be stored in the ICD, while allowing a thinner ICD thickness.
The inventions described in the aforementioned parent patent application (U.S. patent application Ser. No. 09/073,394) advantageously address the above and other needs. In particular, the parent patent application described a technique for generating a highly efficient biphasic defibrillation pulse by switching at least two charged capacitors from a parallel connection to various combinations of a parallel/series connection or a series connection during the first phase of the defibrillation pulse. Such mid-stream parallel/series connection changes of the capacitors and steps up the voltage applied to the cardiac tissue during the first phase. A stepped-up voltage during the first phase, in turn, gives an extra boost to, and thereby forces additional charge (current) into, the cardiac tissue cells, and thereby transfers more charge to the membrane of the excitable cardiac cell than if the capacitors were continuously discharged in series. Phase reversal is timed with the cell membrane reaching its maximum value at the end of the first phase.
Although the technique of the parent application is quite effective, room for improvement remains. In particular, it would be desirable to provide a technique for generating a defibrillation waveform that requires even less shock energy to reach the myocardial defibrillation threshold so that battery power can be saved and device longevity improved, while still providing effective defibrillation. Moreover, it would be desirable to provide a technique for generating a defibrillation waveform, which reduces the total time, required to reach the myocardial defibrillation threshold thereby permitting the patient to be defibrillated more quickly. It is to these ends that aspects of the invention of the present CIP patent application are primarily directed.
In accordance with a first aspect of the invention, increased myocardial voltage is achieved by a defibrillator configured to generate a defibrillation pulse waveform wherein a first (typically positive) phase of the waveform has at least three distinct voltage peaks. In one embodiment, the defibrillator includes a shocking circuit having a set of first, second and third capacitors and switching circuitry for selectively discharging the capacitors so as to generate the defibrillation pulse waveform having the three-peak positive phase. To this end, the switching circuitry generates a first step of the pulse waveform by discharging the capacitors while all three capacitors are connected in parallel, then generates a second step of the pulse waveform by discharging the capacitors while the first and second capacitors are connected in parallel and the third capacitor is connected in series, and finally generates a third step of the pulse waveform by discharging the capacitors while all three capacitors are connected in series.
Preferably, the switching circuitry is configured to discharge the capacitors during the three steps of the pulse waveform for first, second and third time periods selected to maximize the final myocardial voltage within myocardial tissue receiving the pulse waveform. To maximize the final myocardial voltage, the first, second and third time periods are set to:                     d        1        opt            =                                    -                                          τ                m                                            α                1                                              ·          ln                ⁢                  {                                    (                                                τ                  m                                                  τ                  s1                                            )                        ⁢                          xe2x80x83                        ⁢                          (                                                2                  -                                                            α                      2                                                              α                      1                                                                                        1                  -                                                            α                      2                                                              α                      1                                                                                  )                                }                      ;                      d        2        opt            =                                    +                                          τ                m                                            α                2                                              ·          ln                ⁢                  {                                    (                              1                2                            )                        ⁢                          xe2x80x83                        ⁢                          (                                                2                  -                                                            α                      2                                                              α                      1                                                                                        1                  -                                                            α                      2                                                              α                      1                                                                                  )                        ⁢                          xe2x80x83                        ⁢                          (                                                1                  -                                                            α                      3                                                              α                      2                                                                                                            K                    C                                    -                                                            α                      3                                                              α                      2                                                                                  )                                }                      ;    and              d      3      opt        =                            -                                    τ              m                                      α              3                                      ·        ln            ⁢              {                              (                          K              C                        )                    ⁢                      xe2x80x83                    ⁢                      (                                          1                -                                                      α                    3                                                        α                    2                                                                                                K                  C                                -                                                      α                    3                                                        α                    2                                                                        )                          }              ;
wherein
KC=1+(CC)/(CA+CB+CC);
xcex11=1xe2x88x92(xcfx84m/xcfx84s1), xcex12xc2x7=1xe2x88x92(xcfx84m/xcfx84s2), and xcex13=1xe2x88x92(xcfx84m/xcfx84s3);
xcfx84s1=Rsxc2x7Cs1; xcfx84s2=Rsxc2x7Cs2; and xcfx84s3=Rsxc2x7Cs3;
Cs1=CA+CB+CC;
Cs2=[(CA+CB)xc2x7(CC)]/[CA+CB+CC];
            C      s3        =          1      ⁢              /            ⁢              ⌊                              1                          C              A                                +                      1                          C              B                                +                      1                          C              C                                      ⌋              ;
CA, CB, and CC are the capacitances of the first, second and third capacitors, respectively; and
xcfx84m is a predetermined myocardial tissue time constant.
Also, preferably, the capacitances are selected so as to minimize the amount of energy required to be stored within the capacitors while still achieving the maximum final myocardial voltage. To minimize the amount of energy, the capacitances of the first, second and third time capacitors are set to:                     C        A        opt            =              0.6673        ·                  (                                    τ              m                                      R              s                                )                      ;                      C        B        opt            =              0.6673        ·                  (                                    τ              m                                      R              s                                )                      ;          xe2x80x83        ⁢    and              C      C      opt        =          1.5356      ·              (                              τ            m                                R            s                          )              ;
wherein
Rs is a predetermined system resistance.
Using these capacitance values, the optimal first, second and third time periods d1opt, d2opt, and d3opt for use in maximizing the myocardial potential may be simplified to:
d1opt=0.878xc2x7xcfx84m;
d2opt=0.277xc2x7xcfx84m; and
d3opt=0.200xc2x7xcfx84m.
By employing a three capacitor shocking circuit configured as just summarized, the amount of energy required to reach a myocardial defibrillation threshold is less than for one-capacitor or two-capacitor systems, regardless of the total capacitance of the system. Hence, power can be saved, while still providing effective defibrillation. Moreover, the total time required to reach the myocardial defibrillation threshold is less than with one-capacitor or two-capacitor systems, permitting the patient to be defibrillated more quickly. Additionally, the aforementioned three-capacitor system is generally less influenced by variations in underlying parameters and operating conditions and hence is generally more reliable.
In accordance with a second aspect of the invention, a method is provided for making and using an ICD capable of generating the improved defibrillation pulse waveform. In one embodiment, capacitances are selected for capacitors of a shocking circuit for use within the ICD. Optimal time periods for three steps of the positive phase of the defibrillation pulse to be generated by the shocking circuit are then determined, with the time durations being determined so as to maximize a final myocardial voltage within myocardial tissue receiving a pulse waveform from the shocking circuit based on initial capacitor voltages. An ICD is then fabricated having a shocking circuit employing capacitors having the selected capacitances and the shocking circuit is configured to be capable of selectively discharging the capacitors to generate a defibrillation pulse waveform having step durations equal to the optimal time periods. The ICD is implanted within a patient and is activated to detect defibrillation. The shocking circuit is controlled to generate the defibrillation pulse having the three-step positive phase upon detection of fibrillation. In this manner, maximum myocardial voltage can be achieved within myocardial tissue connected to the ICD.